1. Field of the Invention
The invention relates to methods for making an improved electrode for cardiac pacing and defibrillating. Specifically, the methods of the invention produce, electrodes which demonstrate improved capabilities of stimulating and sensing electrically excitable tissues.
2. Description of the Related Art
In the human heart, a small cluster of cells called the sinus node (SN) constitutes the primary natural cardiac pacemaker. The cardiac impulse arising from the SN is transmitted to the atria on the right and left sides of the heart causing the atria to contract. The impulse from the SN is transmitted via pathways in the atria leading to another group of cells, the atrioventricular node, and then via a conduction system comprising the bundle of His, the right and left bundle branches, and the Purkinje fibers, causing the ventricles to contract. This action is repeated in a rhythmic cardiac cycle in which the atrial and ventricular chambers alternately contract and pump, then relax and fill.
The SN is spontaneously rhythmic and is termed the sinus rhythm. Secondary pacemakers in other cardiac tissues tend to be inhibited by the more rapid rate at which impulses are generated by the SN. A number of factors may affect the rate of sinus rhythm. The slower rates (below 60 bpm) are called sinus bradycardia, and the higher rates (between 101 and 160 bpm) are termed sinus tachycardia.
Disruption of the natural pacemaking and propagation system as a result of aging or disease is commonly treated by artificial cardiac pacing. Pacing is a process by which rhythmic electrical discharges are applied to the heart at a desired rate from an implanted artificial pacemaker. In its simplest form, the pacemaker consists of a pulse generator powered by a self-contained battery pack, and a lead including at least one stimulating electrode for delivery of electrical impulses to excitable myocardial tissue in the appropriate chamber(s) in the right side of the patient's heart. Typically, the pulse generator is surgically implanted in a subcutaneous pouch in the patient's chest. In operation, the electrical stimuli are delivered to the excitable cardiac tissue via an electrical circuit that includes the stimulating and reference electrodes, and the body tissue and fluids.
Pacemakers range from the simple fixed rate device that provides pacing with no sensing function, to highly complex devices that provide fully automatic dual chamber pacing and sensing functions. The demand ventricular pacemaker, so termed because it operates only on demand, has been the most widely used type, It senses the patient's natural heart rate and applies stimuli only during periods when that rate falls below the pre-set value. The dual function pacemaker is the latest in a progression toward physiclogic pacing--the mode of artificial pacing that restores cardiac function as much as possible toward natural pacing.
There has also been increasing usage of pacing in the management of tachyarrhythmias. Defibrillation ("DF"), the method employed to terminate fibrillation, involves applying one or more high energy "countershocks" to the heart in an effort to overwhelm the chaotic contractions of individual tissue sections, allow reestablishment of an organized spreading of action potential from cell to cell of the myocardium, and thus restore the synchronized contraction of the mass of tissue. The term "cardioversion" is sometimes used broadly to include DF.
Cardiac output is considerably diminished during an episode of ventricular tachycardia ("VT") because the main pumping chambers of the heart, the ventricles, are only partially filled between the rapid contractions of those chambers. Moreover, VT presents a significant risk of acceleration of the arrhythmia into ventricular fibrillation ("VF"), either spontaneously or in response to treatment of the VT. VF is characterized by rapid, chaotic electrical and mechanical activity of the excitable myocardial tissue. VF manifests an instantaneous cessation of cardiac output as the result of the ineffectual quivering of the ventricles. Unless cardiac output is restored almost immediately after the onset of VF, tissue begins to die for lack of oxygenated blood, and patient death will occur within minutes.
From the factors stated above, it is clear that the principal requirements of pacing and defibrillation, delivery of the pulse and sensing of the electrical state of the target tissue, depend heavily on the abilities of the electrodes. These functions must be routinely and unfailingly carried out over extended device implantation lifetimes. Improvements in electrodes which enhance the pulse-delivering or sensing functions or which reduce power consumption to achieve these ends are needed. It would be especially valuable to be able to achieve a more natural pacing regimen with smaller pulse generators, batteries and electrodes.
The lead assembly of a pacing electrode consists of an electrically conducting wire that is insulated from the tissue. One end of the wire connects to the pulse generator while the other end has an electrode adapted to stimulate excitable myocardial tissue on the inner surface of the heart, the endocardium (an endocardial electrode), or to the outer surface of the heart, the epicardium (an epicardial electrode). A second electrode is also connected to the body at a position through which the electrical circuit is completed in connection with body tissue and fluids. In most cases, an endocardial lead design is used for the implantable cardiac pacemaker because it can readily be inserted through a vein to introduce the stimulating electrode into the chamber to be paced. Epicardial leads require a thoracotomy to affix the stimulating electrode to the heart.
The implantable defibrillator (tachycardia pacer) essentially consists of a pulse generator powered by a combination of batteries and capacitors, and a lead assembly, In this case, the charge delivered to defibrillate the heart is several orders of magnitude larger than that from a cardiac pacemaker. However, in both cases, the stimulation depolarizes a critical mass of the heart.
Typically, electrodes for defibrillation are larger than those used for cardiac pacing because a greater area of the heart tissue needs to be stimulated. These electrodes may be in the form of patches applied directly to the heart. The most common approach in the past has been to suture two patches to the epicardial tissue via thoracotomy. It has been theorized that electrodes with large surface areas are important for a wider distribution of current flow and a more uniform voltage gradient over the ventricles. Others have postulated that uniformity of current density is important since: (i) low gradient areas contribute to the continuation or reinitiation of ventricular fibrillation, and (ii) high current areas may induce temporary damage, that then may cause sensing difficulties, set-up areas of reinitiation of fibrillation, or even potentially cause permanent damage (new arrhythmias, decreased contractility, and myocardial necrosis).
The modern trend in tachycardia pacing has been to use transvenous leads instead of thoracotomy systems. The electrodes in the lead assembly generally consist of two coil electrodes approximately 2 to 4 inches long, one placed in the right ventricle (RV) and the other in the superior vena cava (SVC) or the left brachycephalic vein. The shocking electrodes may be bipolar, in which case only one lead is used, or a combination of two leads, or one endocardial electrode and one subcutaneous patch or epicardial electrode, or two leads and a subcutaneous patch.
Factors that influence the success of defibrillation shocks include the underlying physiclogic substrates of the heart, the pulse waveform of the shock, and the electrode system. In many ways, the physical criteria necessary to design bradycardia electrodes equally apply to tachycardia or defibrillation electrodes.
Two types of lead designs are in common use today, a unipolar (one wire, one electrode) and a bipolar (two wires, two electrodes) lead system. With the unipolar lead, the stimulating electrode is paced against a reference electrode remotely placed from the heart. This reference electrode is usually the metal pacemaker can. For bipolar stimulation, the reference and the stimulating electrodes are normally in close proximity to one another on the same lead and usually in the same chamber. The reference electrode in this case is a ring or sleeve electrode placed a few millimeters from the stimulating tip electrode.
In operation, the pulse generator delivers an output pulse via the lead for electrical stimulation of the excitable myocardial tissue. Stimulation is a function of the current density, i.e., current per unit area. The current required to produce a given current density decreases in direct proportion to the electrode's active or microscopic surface area.
The current from the pacemaker is also affected by a combination of the electrode impedance, the nature of the electrode-endocardial tissue/electrolyte interface, and the impedance of the pacemaker circuitry. Since modern pacemakers operate in a range between 1-2 kHz frequency, the circuit impedance becomes insignificant during pulsing when compared to that due to the electrode impedance and the electrode-endocardial tissue/electrolyte interfacial impedance (commonly termed "spreading" impedance). Hence, the electrode design and materials determine the overall current requirements of the system.
The spreading impedance of an electrode depends predominantly on the tissue resistivity affected by the overall size and shape of the electrode material, the surface characteristics of the electrode, and its reactivity with the tissue. The electrode impedance occurs within a few thousand angstroms up to a few microns from the electrode surface, and results from the charge-transfer reactions taking place at the electrode/electrolyte interface. The electrode impedance is affected by the surface area and nature of the electrode material. The impedance of the output pulse generated by the pacemaker is proportional to the macroscopic geometric surface area of the electrode and the radius of the electrode.
Stimulation requires that an electric field of adequate field strength and current density be imposed on the excitable myocardial tissue in the vicinity of the electrode to initiate rhythmic contractions. The minimum electrical pulse necessary to produce such contractions is referred to as the stimulation threshold. The greater the efficiency of the electrode to generate contractions, the smaller is the amplitude and/or duration of the pulse required to exceed the threshold. The stimulation threshold is affected by the electrode material, electrode geometry, and electrode-tissue interactions. In essence, highly efficient electrodes with low threshold voltages are desirable in order to conserve battery life. It has also been theorized that a high efficiency electrode with a lower voltage threshold and a correspondingly lower energy consumption for tissue stimulation reduces injury to tissue at the stimulation site.
At the time of implant, the acute stimulation threshold is two to three times lower than the chronic stimulation threshold observed later. The increase in threshold is attributed to a fibrous capsule which develops around the electrode tip, i.e., the development of a layer or layers of unexcitable connective tissue surrounding the electrode tip at the stimulation site. The fibrotic growth results in a virtual electrode surface area which is considerably greater than the actual surface area of the electrode. This increase in surface area lowers current densities at the tip and results in a higher stimulation threshold. The thickness of the fibrous capsule around the electrode tip is generally dependent on the fixation characteristics at the time of implant, the geometry of the electrode tip, the microstructure of the electrode tip, and the material used for the electrode. It may also be dependent on the current density at the electrode/tissue interface during the pacing pulses. A lower current density may result in less myocardial damage and hence, lead to a thinner fibrous capsule around the electrode tip. On the contrary, electrodes that have a rough surface microstructure or have sharp protrusions may be too abrasive, thereby causing irritation leading to the development of a thicker fibrous capsule.
In addition to pacing functions, the electrode must function to sense the activity of the heart by determining the aberrant behavior in the ventricular rhythms so that pacing operation will be initiated. The frequencies at which signals are typically sensed are in the bandwidth of 20-100 Hz. In these frequencies, the electrode-endocardial tissue/electrolyte interfacial impedance becomes significant. Interfacial impedance is affected by the microscopic surface area of the electrode and is established within a few microns of the electrode's surface. The microscopic surface area of the electrode is represented by all wettable surfaces including interstitial porosity, surface cracks, crevices, and channels on the surface of the stimulating electrode. Electrodes with a higher intrinsic surface area are desirable for greater sensing of the heart's activity.
Depending on the applied potential and pulse duration, activities at the electrode interface generally involve charge transfer across the electrode-tissue/electrolyte interface by a combination of faradaic processes or oxidation-reduction reactions and double layer charging. As current densities increase, these reactions change the ionic concentration at the interface, requiring migration of ions from increasingly greater distances. The greater the current density, the larger are the polarization losses on the electrode. The concentration gradient set-up at the electrode/electrolyte interface is the source of the after potential.
Current density is related to the pacing threshold and sensing capability (amplitude of the depolarization events), i.e., if the current density is too high, the electrode is perturbed more from its initial equilibrium voltage thereby decreasing its sensing capabilities. If the current density is low, the voltage of the electrode is less perturbed and therefore sensing is less affected. Sensing is at its most optimum at a lower current density. However, a finite current density is required for cardiac muscle depolarization. Certain improvements in sensing have been achieved (see; e.g., U.S. Pat. No. 5,267,564 which relates to a pacemaker lead for sensing a physiological parameter of the body, a portion of which lead comprises a platinum-iridium outer cap).
In all types of stimulation electrodes, the electrode itself must be both chemically corrosion resistant and mechanically stable enough to withstand chronic application. It must possess a high charge capacity. It must also inject a substantial level of electric charge into the tissue to be stimulated. Finally, the ability to inject charge must not deteriorate significantly over time after implantation.
Stimulation of tissues requires that the charge be injected reversibly by a purely capacitive mechanism. In such a mechanism, the electrode behaves as a charge flow transducer between media exhibiting different charge flow properties. The capacitive mechanism allows electrons to flow away from the stimulation electrode causing electrical charges at the electrode/electrolyte interface to orient themselves in order to cause a displacement current through the electrolyte. Since the electrolyte is an ionic medium, the slight displacement of the ions in reorientation creates a charge flow.
When irreversible chemical reactions begin to occur, the mechanism is no longer capacitative. Irreversible faradaic reactions may lead to water electrolysis, oxidation of soluble species, and metal dissolution. In addition, some of the products of the reactions may be toxic. Neither gas evolution nor oxide formation reactions contribute to electrical stimulation of excitable tissue. The stimulation energy is wasted in electrolyzing the aqueous phase of blood instead of carrying desirable charged species from one electrode to the other via the tissues. Stimulation electrodes should preferably allow a large charge flow across the electrode-tissue interface without the risk of irreversible faradaic reactions. Selection of the metal of the electrode is critical.
A metal of choice in electrode manufacturing has traditionally been titanium. On a fresh titanium surface, however, oxygen ions react with the titanium anode to form an oxide layer. Once a finite oxide thickness has been formed on the surface, polarization increases further. A point is reached when the oxygen ions reaching the surface of the titanium cannot be reduced further to form the oxide, and instead are reduced to elemental oxygen to form oxygen gas. The oxide film developed on the surface of a titanium electrode, either naturally or electrochemically, is irreversible. It cannot be reduced to the original metal by passing a charge in the reverse direction. Hence, it is clear that virgin titanium metal is a poor choice for electrode construction since it forms a semi-conductive oxide on its surface before and even during the electrical stimulation. Platinum, and much more so stainless steel, have been shown to undergo irreversible dissolution during stimulation as well.
Titanium oxidation reactions are several times more likely in an oxidative environment than those of platinum or platinum alloys, but a thousand times less so than those of stainless steel. Unfortunately, due to the expense of platinum metal and the requirement for large amounts of metal in patch-type electrodes, production costs are too high for platinum electrodes. Therefore, even though oxidation problems are more prevalent in them, titanium electrodes are typically used.
From the equation C=k.epsilon.A/d: where .epsilon. is the permittivity of vacuum, A is the real surface area of the film, k is the dielectric constant of the film, and d is the thickness of the porous material, it can be seen that in order to achieve a large charge-storage capacity, the porosity of the dielectric may be maximized with a large accessible surface area. Numerous types of cardiac pacing and defibrillation electrodes have heretofore been developed with these and other factors in mind, utilizing various configurations and materials asserted to promote lower stimulation thresholds and improved electrical efficiencies. Thus, for both bradycardia and tachycardia applications, it is desirable to minimize the electrical impedance at the electrode-tissue interface by increasing the intrinsic surface area of the electrode or by reducing formation of a capsule of inactive tissue that surrounds and isolates the electrode from active tissue,
Microporous electrodes based on sintered titanium, sintered titanium nitride, and microporous carbon or graphite have been used with some degree of success. However, the electrode reactions in aqueous solutions involve significant gas generation similar to the behavior of native titanium. Sanding or sandblasting electrode surfaces is a broadly used method to achieve surface area enhancement. For example, French Patent 2,235,666 relates to a stainless steel electrode tip which is sanded to increase surface area and reduce the impedance of the electrode.
Other methods have also been used. U.S. Pat. No. 5,318,572 relates to a platinum-iridium (90:10) porous electrode with recess slots in the shape of a cross and at least one, preferably two variably-sized, porous coating/s of 20-80 micron diameter platinum-iridium (90:10) spheres deposited on the surface of the electrode. On top of this structure, a reactively sputtered coating of titanium nitride is applied. U.S. Pat. No. 4,156,429 describes a means for increasing the reactive surface area by forming a highly porous sintered electrode body consisting of a bundle of fibers, preferably of platinum but alternatively of Elgiloy, titanium, or a platinum-iridium alloy. Conversely, the fibers may be encompassed within a metallic mesh to yield a seventy percent to ninety-seven percent porosity, U.S. Pat. No. 5,203,348 relates to defibrillation electrodes which can be formed on titanium ribbons or wires with a sputtered outer layer of platinum, or a silver core in a stainless steel tube with a platinum layer formed onto the tube. A divisional of that patent (U.S. Pat. No. 5,230,337) discloses that the coating is preferably made by sputtering to apply a "microtexture" to increase the surface area of the electrode.
U.S. Pat. No. 5,178,957 relates to electrodes and a method of making electrodes including pretreatment of the surface by sputter-etching and sputter-depositing a noble metal on the surface. U.S. Pat. No. 5,074,313 relates to a porous electrode with an enhanced reactive surface wherein surface irregularities are introduced to increase surface area by glow discharge or vapor deposition upon sintered wires. U.S. Pat. No. 4,542,752 describes a platinum or titanium substrate coated with a porous sintered titanium alloy which in turn is coated with a porous carbon. The latter was claimed to promote tissue ingrowth and provide low polarization impedance. U.S. Pat. No. 4,784,161 relates to making a porous pacemaker electrode tip using a porous substrate, where the porous substrate is preferably a non-conductive material such as a ceramic or a polymer made porous by laser drilling, sintering, foaming, etc. to result in pores 5-300 microns in depth. U.S. Pat. No. 4,603,704 features a hemispherical electrode made of platinum or titanium, coated with a porous layer consisting of a carbide, nitride, or a carbonitride of at least one of the following metals; titanium, zirconium, hafnium, molybdenum, niobium, vanadium, or tungsten. U.S. Pat. No. 4,281,668 discloses a vitreous carbon or pyrolytic carbon electrode that is superficially activated, e.g., by oxidation, for microporosity. The electrode is then coated with a biocompatible ion-conducting, hydrophobic plastic. The latter is said to substantially prevent thrombus formation.
Despite the numerous means of increasing the surface area to reduce polarization losses and after potentials and the use of noble metals and their alloys as electrodes as described above, with varying degrees of success, there remain significant problems pertaining to polarization losses and sensing difficulties. In order to make further improvements to the electrode, stable oxides of some of these noble metals have been employed as a coating,
It is known that certain metals, metallic oxides, and alloys are stable during electrolysis, and that these metals are useful in a variety of electrode applications such as chlor-alkali electrolysis (see, e.g., U.S. Pat. No. 5,298,280). Such metals typically include the members of the platinum group; namely, ruthenium, rhodium, palladium, osmium, iridium, and platinum. These metals are not suitable for construction of the entire electrode since their cost is prohibitive. Therefore, these metals or their alloys, or as metallic oxides, have been applied as a thin layer over a strength or support member such as a base member made of one of the valve metals (Ti, Ta, Nb, Hf, Zr, and W). These valve metals or film-forming metals as they are sometimes known, are much less expensive than platinum group metals and they have properties which render them corrosion resistant. However, as previously mentioned, they lack in good surface electroconductivity because of their tendency to form on their surface an oxide having poor electroconductivity.
As noted previously, titanium is generally the metal substrate of choice since it is lightweight and relatively inexpensive compared to the other metallic substrates. However, Ti has a naturally occurring oxide passivated on its surface having a rutile structure. This oxide is fairly non-conductive and has to be removed before titanium can fully function as an electrically conductive substrate. Various procedures have been employed in prior art to "etch" this film away. For instance, U.S. Pat. No. 5,181,526 relates to an electrode comprising platinum or titanium and a mixture of platinum and a platinum group metal oxide coated thereon, where an upper portion of the electrode is a mesh or is porous, and the electrode head is electrolytically corroded to remove the oxide using NaCl-HCl or hot oxalic acid solution prior to deposition of the platinum-iridium coating.
It is known that titanium oxide and the oxides of the other valve metals have better semi-conducting properties than the native oxide when doped with other elements or compounds which disturb the lattice structure and change the conductivity of the surface oxide. Metal oxides other than titanium oxide when intimately mixed and heated together have the property of forming semiconductors, particularly mixed oxides of metals belonging to adjacent groups in the Periodic Table. It is also known that platinum group metals and platinum group metal oxides may be coated on the surface of the valve metals to achieve this semi-conducting properties. U.S. Pat. No. 4,717,581 teaches the use of iridium oxide coated electrodes for neural stimulation. A metallic electrode made of platinum, platinum-iridium alloy, stainless steel, stainless steel alloys, titanium, titanium alloys, tantalum, and tantalum alloys is coated with iridium oxide to form the electrode. U.S. Pat. No. 4,679,572 discloses an electrode with a conductive tip portion and a substrate composed of a material conventionally employed for pacing electrodes, and a layer of film of iridium oxide overlying the surface of the substrate. The tip portion may be provided with recesses to which the iridium oxide surface layer is confined.
Valve metals have the capacity to conduct current in the anodic direction and to resist the passage of current in the cathodic direction, (i.e., the anodic reaction is irreversible) and are sufficiently resistant to the electrolyte media. In the anodic direction, however, their resistance to the passage of current goes up rapidly, due to the formation of an oxide layer thereon, so that it is no longer possible to conduct current to the electrolyte in any substantial amount without substantial increases in voltage which makes continued use of uncoated valve metal anodes in the electrolytic process uneconomical and inefficient.
In order to avoid this passivation on the surface of the valve metal, a metal oxide or a mixed metal oxide of the platinum group is used. The oxide from this group is very stable and does not grow further. In addition, it provides a protection for the underlying metal. Many of these oxides are generally reversible to aqueous based redox species and hence undergo reversible redox reactions with species such as hydrogen ions and hydroxyl ions leading to the formation of higher oxidation state surface oxides.
Electrodes capable of more natural pacing and defibrillation are needed. Improved electrodes should have the following features for efficient stimulation of the myocardial tissue: smaller geometric macroscopic surface area and smaller electrode radius; higher intrinsic microscopic surface area; and appropriate surface nature, for achieving: (1) finite low current drain; (2) finite current density; (3) high pacing impedance; (4) low sensing impedance; (5) greater efficiency to produce contractions of the heart wall at lower voltage threshold; and (6) lower tissue irritations and hence lesser fibrotic growth.